Conventional treatment of bone defects requires the use of either non-resorbable or resorbable prosthetic structures. The resorbable structures or materials either support the in growth of adjacent bone and soft tissue or actively induce the formation of new bone. The active formation of new bone, termed osteoinduction, occurs only in the presence of demineralized bone matrix or in the presence of protein extracts from such matrix (biomaterials), or a combination of both materials, preferably with interconnected porous spaces across the substratum of the biomaterial. This allows bone growth into the porous spaces of the biomaterial, securing its incorporation and osteointegration with the surrounding viable bone at the margins of the bone defect. Such porous biomaterials, which allow bone growth into their porous spaces, are defined as osteoconductive biomaterials.
The necessity of having viable bone in direct contact with the porous biomaterial to ensure adequate bone ingrowth via osteoconduction is, however, a limiting factor particularly in large bony defects, since the depth of bone penetration within the porous spaces may be confined to the peripheral regions of the implant only. Thus, osteointegration often does not occur or is not maintained along the entire implant surface. One approach for preparing an osteoinductive material has been to adsorb onto its surfaces exogenous growth and morphogenetic factors (collectively, bone morphogenetic proteins (BMPs)), which are capable of inducing differentiation of bone within the porous spaces of the biomaterial. See e.g., Jefferies et al., in U.S. Pat. Nos. 4,394,370; 4,472,840; 6,311,690 describing bone graft materials comprising collagen and demineralized bone matrix or extracted BMP. However, BMPs have limited shelf life and risk of adverse systemic effects. Therefore, a preferred alternative would be a bioactive material, which is capable of spontaneously initiating bone formation within the porous spaces independent of the presence of viable bone at its interfaces.
“Bioactivity” is a unique property associated with the ability of a synthetic material to interact or bond with living tissue. All materials implanted in vivo elicit a response from the surrounding tissue. Four types of response are possible: (i) tissue death if the material is toxic; (ii) replacement by the surrounding tissue if the material is nontoxic and dissolves; (iii) formation of a fibrous tissue capsule of variable thickness if the material is nontoxic and biologically inactive; and (iv) formation of an interfacial bone if the material is nontoxic and biologically active. Bioactive materials fall into the fourth category.
Hydroxyapatite is the most recognized bioactive material for bone tissue. The presence and formation of calcium hydroxyapatite at the implant-bone interface appears to be critical for bone bonding, and it is one of the key features necessary for successful bioactive bone implants. Calcium hydroxyapatite coatings on implants or calcium hydroxyapatite blocks have been used to produce implants with bone-binding abilities, e.g., U.S. Pat. No. 6,302,913 (Ripamonti) introduced the concept that the shape and configuration of the hydroxyapatite implant regulates the initiation of bone formation in vivo.
Moreover, it has been shown that several crystalline and amorphous (glassy) oxide materials can induce the growth of hydroxyapatite in the environment of simulated body fluid. Among them is silica (SiO2), but several other oxides, such as TiO2, ZrO2 and Ta2O5, and various silica-based glasses are also effective. For example, through the use of an in vitro immersion method using a simulated physiological solution that mimics the ion concentration found in body fluids, the formation of the calcium hydroxyapatite layers on bioactive glasses, bioactive glass-ceramics and polymers have been produced. As a result, this method of “in vitro immersion” has been used to predict bone-bonding potential of bone implant materials (Kokubo et al., J. Biomed. Mater. Res. 24:721-734 (1990); Li et al., J. Biomed. Mater. Res. 34:79-86 (1997)). Therefore, these materials can also be regarded as bioactive. For example, at high pH, on the SiO2 surface, the glass-solution interface is crystallized into a mixed hydroxyapatite phase of the CaO and P2CO5 that is released into solution during the network dissolution. Then, the hydroxyapatite crystallites nucleate and bond to the interfacial biological metabolites, such as mucopolysaccharides, collagen and glycoproteins.
Currently, bioactive powders are produced by conventional processing techniques well-known in the art. The various constituents (e.g., reagent-grade Na2CO3, CaCO3, P2CO5 and SiO3) are usually mixed in a suitable mixing device such as a rolling nill, and then heated to a temperature (generally 1250-1400° C.) sufficient to cause the particles to melt and coalesce. See, e.g., U.S. Pat. No. 4,775,646; Ogino et al., J. Biomed. Mat. Res. 14:55-56 (1980). However, the use of such high temperatures and specialized equipment results in significant production costs. Moreover, conventional bioactive glass compositions tend to require an alkali metal oxide such as Na2O to serve as a flux or aid in melting or homogenization, resulting in a high pH at the interface between the glass and surrounding fluid or tissue. Unfortunately in vivo, this can induce inflammation. Moreover, the rate of tissue repair, which drives the interfacial tissue-glass bonding promoted by bioactive material, tends to vary within a narrow pH range, and if the surrounding environment is too acidic or alkaline, repair shuts down, and interfacial bonding is defeated. Consequently, high rates of bioactivity (as measured by surface hydroxyapatite accretion) tend to be associated with significant local pH changes due to the release of alkali metal oxide ions.
Conventional glasses also tend to be difficult to mix to homogeneity, making quality control a problem for materials intended for implantation in the body. This is due to the relatively large grain size of the glass precursors, which generally measure approximately 10 to 1000 microns in diameter. It is difficult to obtain “molecular scale” mixing, i.e., homogeneity at the molecular level, using ordinary mixing techniques, such as stirring of the relatively viscous silicate melts.
Currently bioactive powders are limited to a SiO2 content that is less than 60 mole %, which is problematic because, for example, the rate of hydroxyapatite formation is dependent upon SiO2 content. Therefore, compatibility between the bioactive material and the surrounding tissue is maximized when the material's bioactivity rate (i.e., the speed at which hydroxyapatite is produced) matches the body's metabolic repair rate. However, an individual's repair rate can vary with age and disease state, among other factors, rendering identification of a single, ideal bioactivity rate impossible.
The SiO2 level also determines the thermal expansion coefficient and elastic modulus of the glass. Particularly in the case of porous compositions, the ability to coat the glass onto a strong substrate (e.g., metal) significantly increases the range of clinical applications to which the glass will be amenable. Such coating is most conveniently accomplished when the thermal expansion coefficient of the glass matches that of the substrate, but restrictions on SiO2 variation diminish the available range of coefficients. Particular values or ranges for the elastic modulus can also be important in certain clinical applications (such as avoiding stress shielding of the repair of long bones and joints). Consequently, some glass compositions are unsuitable if the SiO2 level cannot be adjusted.
Calcium phosphate-based ceramics and glasses have the ability to bond with bone tissues and have been widely used in bone repair (Gross et al., Ann. NY Acad. Sci. (ed. Ducheyne and Lemons) 523 (1988); U.S. Pat. No. 6,328,990). Based on a comparison of literature data, it was suggested that 45S5 “bioactive” glass (45% SiO2, 24.5% Na2O, 24.5% CaO, and 6% P2O5) had the highest rate of bonding to bone (Hench, Ann. NY Acad. Sci. (ed. Ducheyne and Lemons) 523:54-71 (1988)). Recently, 45S5 bioactive glass has been considered for use as bioactive ceramic microspheres in 3D bone cell cultures in rotating bioreactors (Qiu et al., Tissue Engineer 4:19-34 (1998)). The use of bone bioactive materials is of great interest in bone synthesis in vitro because of their ability to promote cell-material bonding and the potential to enhance bone formation.
Solid bioceramic microspheres typically have a density higher than 2 g/cm3. When used in bioreactors, the solid ceramic microspheres experience a high shear stress, which causes cell detachment and damage (Qiu et al., 1998). One way of solving this problem has been to reduce the apparent density of the microspheres through a hollow structure approach (Qiu et al., Biomaterials 20:989-1001 (1999)). Cell culture studies have confirmed that the hollow bioceramic microspheres (SiO2/Al2O3/CaP) experience a low shear stress and can support 3D bone cell cultures in rotating bioreactors. However, because of their non-degradable component, Al2O3, hollow bioceramic microspheres cannot be completely replaced by bone tissues.
Qi et al., (Chem. Mater. 10:1623-1626 (1998)), described the formation of mesoporous silica spheres by a process using a cationic-nonionic surfactant mixture in aqueous acidic conditions. A typical synthesis involved stirring an aqueous acidic solution of a cationic ammonium surfactant, and a nonionic surfactant (decaethylene glycol monohexadecylether), to which an alkoxysilane was added (TEOS). The resulting material has a high surface area (1042 m2/g) and ˜5 μm particle size, and could be used for use as a chromatographic matrix. However, to utilize these mesoporous spheres for tissue engineering applications, they would first have to be assembled into a form that could be easily handled, and this has not been demonstrated. In addition, their bioactivity is unknown, and a long synthesis time (16 hours) would be required, making the use of such a mixture of surfactants impractical for commercial use.
Other processes have been developed for synthesizing mesoporous silica spheres in acidic solution (e.g., U.S. Pat. No. 6,334,988 (Gallis et al.), Ozin et al., J. Mater. Chem. 8(3):743-750 (1998)); Schacht et al., Science 273:768-771 (1996)) describing an emulsion process for synthesizing mesoporous silica spheres). A silicon alkoxide (TEOS) was dissolved in an organic solvent, typically mesitylene, and the mixture was slowly added to an aqueous acidic solution containing a cationic ammonium surfactant (CTAB). Schacht found that by varying the stir rate during the course of the reaction, the particle morphology could be changed. At slower stirring rates, the reaction mixture produced microspheres and some transient solid fibers; however, as the stirring rate was increased, the amount of fibers decreased with the increasing amounts of spheres. Scanning Electron Microscopy (SEM) indicated the final particles were hollow and spherical in nature. It was shown that these hollow spheres were brittle, and could be crushed with a spatula. However, the brittle nature of the spheres, in combination with the fact that they were not porous throughout their interior, were unfavorable characteristics for use as a chromatographic matrix.
The foregoing processes produce materials which exhibit regular powder X-ray diffraction patterns with one or more relatively narrow diffraction peaks, indicating that they contain a relatively ordered arrangement of pores, and are similar to SBA-3, a mesoporous material with a hexagonal arrangement of linear pores (Huo et al., Science 268:1324(1995)). SBA-3 is similar to the more widely known MCM-41, which has an identical arrangement of pores, but is synthesized in basic solution (Kresge et al., Nature 359:710-712 (1992)). While mesoporous silica having such ordered pores has use in a variety of contexts, they have not been explored for tissue engineering applications.
In an alternative process, U.S. Pat. No. 5,074,916 (Hench) teaches the use of sol-gel technology to synthesize bioactive glass powders from SiO2—CaO—P2O5. The production of ceramic and glass materials by the sol-gel process has been known for many years. A “sol” is a dispersion of colloidal particles in a liquid, while a “gel” denotes an interconnected, rigid network with pores of submicrometer dimensions and polymeric chains, having an average length >1 μm. These pores are typically filled with air, so that the gel is sometimes referred to as an aerogel. Basically, the sol-gel process involves mixing of the glass precursors into a sol; casting the mixture in a mold; gelation of the mixture, whereby the colloidal particles link together to become a porous three-dimensional network; aging of the gel to increase its strength; drying the liquid from the interconnected pore network; dehydration or chemical stabilization of the pore network; and densification, to produce structures with ranges of physical properties (e.g., Hench et al., Chem. Rev. 90:33 (1990)). However, while such gels have nanometer-sized and mesoporous pores, they lack larger, macroscopic pores equivalent to those that exist in natural bones. Moreover, like the mesoporous silica spheres, the sol-gels tend to be brittle and mechanically weak.
Thus, a new bioactive and biodegradable composite material has, until the present invention, been needed that could be easily handled, without the brittleness and incompatibility of the prior art bone replacement materials, and that combined merits of nanopore materials (such as sol gel glass) and a fibrous construct, without the limitations of solid fibers. Although the existing fibers in the art appear to be “transparent” to fluid, they are nonporous, and therefore lack high specific surface area because they lack a high percentage of nanopores. Accordingly, optimally such material would be characterized by a combination of nanopores for diffusion and delivery, as well as a network of bone like mesopores and macropores to enhance osteoinduction when the material is used in 3-dimensional bone tissue engineering and bone implant materials.